The present invention pertains generally to implantable medical devices and, more particularly, to implantable medical devices fabricated of self-supporting laminated films fashioned into geometric configurations adapted to specific medical uses. More particularly, the present invention relates to metal films, foils, wires and seamless tubes, with increased mechanical properties, which are suitable for use in fabricating implantable endoluminal grafts, stent-grafts and stent-graft-type devices. More specifically, the present invention comprises endoluminal grafts, stent-grafts and stent-graft-type devices that are fabricated entirely of self-supporting laminated films, foils, wires or seamless tubes made of biocompatible metals or of biocompatible materials which exhibit biological response and material characteristics substantially the same as biocompatible metals, such as for example composite materials.
As opposed to wrought materials that are made of a single metal or alloy, these inventive materials are made of at least two layers formed upon one another into a self-supporting laminate structure. Laminate structures are generally known to increase the mechanical strength of sheet materials, such as wood or paper products. Laminates are used in the field of thin film fabrication also to increase the mechanical properties of the thin film, specifically hardness and toughness. Laminate metal foils have not been used or developed because the standard metal forming technologies, such as rolling and extrusion, for example, do not readily lend themselves to producing laminate structures. Vacuum deposition technologies can be developed to yield laminate metal structures with improved mechanical properties. In addition, laminate structures can be designed to provide special qualities by including layers that have special properties such as superelasticity, shape memory, radio-opacity, corrosion resistance etc.
Metal foils, wires and thin-walled seamless tubes are typically produced from ingots in a series of hot or cold forming steps that include some combination of rolling, pulling, extrusion and other similar processes. Each of these processing steps is accompanied by auxiliary steps that include cleaning the surfaces of the material of foreign material residues deposited on the material by the tooling and lubricants used in the metal forming processes. Additionally, chemical interaction with tooling and lubricant materials and ambient gases also introduces contaminants. Some residue will still usually remain on the surface of the formed material, and there is a high probability that these contaminating residues become incorporated during subsequent processing steps into the bulk of the wrought metal product. With decreasing material product size, the significance of such contaminating impurities increases. Specifically, a greater number of process steps, and, therefore, a greater probability for introducing contaminants, are required to produce smaller product sizes. Moreover, with decreasing product size, the relative size of non-metal or other foreign inclusions becomes larger. This effect is particularly important for material thicknesses that are comparable to the grain or inclusion size. For example, austenitic stainless steels have typical grain sizes on the order of magnitude of 10-100 micrometer. When a wire or foil with a thickness in this range is produced, there is significant probability that some grain boundaries or defects will extend across a large portion or even across the total thickness of the product. Such products will have locally diminished mechanical and corrosion resistance properties. While corrosion resistance is remedied by surface treatments such as electropolishing, the mechanical properties are more difficult to control.
The mechanical properties of metals depend significantly on their microstructure. The forming and shaping processes used to fabricate metal foils, wires and thin-walled seamless tubes involves heavy deformation of a bulk material, which results in a heavily strained and deformed grain structure. Even though annealing treatments may partially alleviate the grain deformation, it is typically impossible to revert to well-rounded grain structure and a large range of grain sizes is a common result. The end result of conventional forming and shaping processes, coupled with annealing, typically results in non-uniform grain structure and less favorable mechanical properties in smaller sized wrought metal products. It is possible, therefore, to produce high quality homogeneous materials for special purposes, such as micromechanical devices and medical devices, using vacuum deposition technologies.
In vacuum deposition technologies, materials are formed directly in the desired geometry, e.g., planar, tubular, etc. The common principle of the vacuum deposition processes is to take a material in a minimally processed form, such as pellets or thick foils (the source material) and atomize them. Atomization may be carried out using heat, as is the case in physical vapor deposition, or using the effect of collisional processes, as in the case of sputter deposition, for example. In some forms of deposition, a process, such as laser ablation, which creates microparticles that typically consist of one or more atoms, may replace atomization; the number of atoms per particle may be in the thousands or more. The atoms or particles of the source material are then deposited on a substrate or mandrel to directly form the desired object. In other deposition methodologies, chemical reactions between ambient gas introduced into the vacuum chamber, i.e., the gas source, and the deposited atoms and/or particles are part of the deposition process. The deposited material includes compound species that are formed due to the reaction of the solid source and the gas source, such as in the case of chemical vapor deposition. In most cases, the deposited material is then either partially or completely removed from the substrate, to form the desired product.
The rate of film growth is a significant parameter of vacuum deposition processes. In order to deposit materials that can be compared in functionality with wrought metal products, deposition rates in excess of 1 micrometers/hour are a must and indeed rates as high as 100 micrometers per hour are desirable. These are high deposition rates and it is known that at such rates the deposits always have a columnar structure. Depending on other deposition parameters, and most importantly on the substrate temperature, the columns may be amorphous or crystalline but at such high deposition rates microcrystalline structure development can be expected at best. The difficulty is that the columns provide a mechanically weak structure in which crack propagation can occur uninhibited across the whole thickness of the deposit.
A special advantage of vacuum deposition technologies is that it is possible to deposit layered materials and thus films possessing exceptional qualities may be produced (c.f., H. Holleck, V. Schier: “Multilayer PVD coatings for wear protection”, Surface and Coatings Technology, Vol. 76-77 (1995) pp. 328-336). Layered materials, such as superstructures or multilayers, are commonly deposited to take advantage of some chemical, electronic, or optical property of the material as a coating; a common example is an antireflective coating on an optical lens.
It has not been recognized until relatively recently that multilayer coatings may have improved mechanical properties compared with similar coatings made of a single layer. The improved mechanical properties may be due to the ability of the interface between the layers to relieve stress. This stress relief occurs if the interface provides a slide plane, is plastic, or may delaminate locally. This property of multilayer films has been recognized in regard with their hardness but this recognition has not been translated to other mechanical properties that are significant for metal products that may be used in application where they replace wrought metal parts.
A technological step that interrupts the film growth results in discontinuous columns and prevents crack propagation across the entire film thickness. In this sense, it is not necessary that the structure consist of a multiplicity of chemically distinct layers, as it is common in the case of thin film technology where multilayers are used. Such chemical differences may be useful and may contribute to improved properties of the materials.
The surface of a solid, homogeneous material can be conceptualized as having unsaturated inter-atomic and intermolecular bonds forming a reactive plane ready to interact with the environment. In practice, a perfectly clean surface is unattainable because of immediate adsorption of airborne species, upon exposure to ambient air, of O, O2, CO2, SO2, NO, hydrocarbons and other more complex reactive molecules. Reaction with oxygen implies the formation of oxides on a metal surface, a self-limiting process, known as passivation. An oxidized surface is also reactive with air, by adsorbing simple, organic airborne compounds. Assuming the existence of bulk material of homogeneous subsurface and surface composition, oxygen and hydrocarbons may adsorb homogeneously. Therefore, further exposure to another environment, such as the vascular compartment, may be followed by a uniform biological response.
Current metallic vascular devices, such as stents, are made from bulk metals made by conventional methods which employ many steps that introduce processing aides to the metals make stent precursors, such as hypotubes. For example, olefins trapped by cold drawing and transformed into carbides or elemental carbon deposit by heat treatment, typically yield large carbon rich areas in 316L stainless steel tubing manufactured by cold drawing process. The conventional stents have marked surface and subsurface heterogeneity resulting from manufacturing processes (friction material transfer from tooling, inclusion of lubricants, chemical segregation from heat treatments). This results in formation of surface and subsurface inclusions with chemical composition and, therefore, reactivity different from the bulk material. Oxidation, organic contamination, water and electrolytic interaction, protein adsorption and cellular interaction may, therefore, be altered on the surface of such inclusion spots. Unpredictable distributions of inclusions such as those mentioned above provide unpredictable and uncontrolled heterogeneous surface available for interaction with plasma proteins and cells. Specifically, these inclusions interrupt the regular distribution pattern of surface free energy and electrostatic charges on the metal surface that determine the nature and extent of plasma protein interaction. Plasma proteins deposit nonspecifically on surfaces according to their relative affinity for polar or non-polar areas and their concentration in blood. A replacement process known as the Vroman effect, Vroman L., The importance of surfaces in contact phase reactions, Seminars of Thrombosis and Hemostasis 1987; 13(1): 79-85, determines a time-dependent sequential replacement of predominant proteins at an artificial surface, starting with albumin, following with IgG, fibrinogen and ending with high molecular weight kininogen. Despite this variability in surface adsorption specificity, some of the adsorbed proteins have receptors available for cell attachment and therefore constitute adhesive sites. Examples are: fibrinogen glycoprotein receptor IIbIIIa for platelets and fibronectin RGD sequence for many blood activated cells. Since the coverage of an artificial surface with endothelial cells is a favorable end-point in the healing process, favoring endothelialization in device design is desirable in implantable vascular device manufacturing.
Normally, endothelial cells (EC) migrate and proliferate to cover denuded areas until confluence is achieved. Migration, quantitatively more important than proliferation, proceeds under normal blood flow roughly at a rate of 25 μm/hr or 2.5 times the diameter of an EC, which is nominally 10 μm. EC migrate by a rolling motion of the cell membrane, coordinated by a complex system of intracellular filaments attached to clusters of cell membrane integrin receptors, specifically focal contact points. The integrins within the focal contact sites are expressed according to complex signaling mechanisms and eventually couple to specific amino acid sequences in substrate adhesion molecules (such as RGD, mentioned above). An EC has roughly 16-22% of its cell surface represented by integrin clusters. Davies, P. F., Robotewskyi A., Griem M. L. Endothelial cell adhesion in real time. J. Clin. Invest. 1993; 91:2640-2652, Davies, P. F., Robotewski, A., Griem, M. L., Qualitative studies of endothelial cell adhesion, J. Clin. Invest. 1994; 93:2031-2038. This is a dynamic process, which implies more than 50% remodeling in 30 minutes. The focal adhesion contacts vary in size and distribution, but 80% of them measure less than 6 μm2, with the majority of them being about 1 μm2, and tend to elongate in the direction of flow and concentrate at leading edges of the cell. Although the process of recognition and signaling to determine specific attachment receptor response to attachment sites is incompletely understood, regular availability of attachment sites, more likely than not, would favorably influence attachment and migration. Irregular or unpredictable distribution of attachment sites, that might occur as a result of various inclusions, with spacing equal or smaller to one whole cell length, is likely to determine alternating hostile and favorable attachment conditions along the path of a migrating cell. These conditions may vary from optimal attachment force and migration speed to insufficient holding strength to sustain attachment, resulting in cell slough under arterial flow conditions. Due to present manufacturing processes, current implantable vascular devices exhibit such variability in surface composition as determined by surface sensitive techniques such as atomic force microscopy, X-ray photoelectron spectroscopy and time-of-flight secondary ion mass spectroscopy.
There have been numerous attempts to increase endothelialization of implanted stents, including covering the stent with a polymeric material (U.S. Pat. No. 5,897,911), imparting a diamond-like carbon coating onto the stent (U.S. Pat. No. 5,725,573), covalently binding hydrophobic moieties to a heparin molecule (U.S. Pat. No. 5,955,588), coating a stent with a layer of blue to black zirconium oxide or zirconium nitride (U.S. Pat. No. 5,649,951), coating a stent with a layer of turbostratic carbon (U.S. Pat. No. 5,387,247), coating the tissue-contacting surface of a stent with a thin layer of a Group VB metal (U.S. Pat. No. 5,607,463), imparting a porous coating of titanium or of a titanium alloy, such as Ti—Nb—Zr alloy, onto the surface of a stent (U.S. Pat. No. 5,690,670), coating the stent, under ultrasonic conditions, with a synthetic or biological, active or inactive agent, such as heparin, endothelium derived growth factor, vascular growth factors, silicone, polyurethane, or polytetrafluoroethylene, U.S. Pat. No. 5,891,507), coating a stent with a silane compound with vinyl functionality, then forming a graft polymer by polymerization with the vinyl groups of the silane compound (U.S. Pat. No. 5,782,908), grafting monomers, oligomers or polymers onto the surface of a stent using infrared radiation, microwave radiation or high voltage polymerization to impart the property of the monomer, oligomer or polymer to the stent (U.S. Pat. No. 5,932,299).
Thus, the problems of thrombogenicity and re-endothelialization associated with stents have been addressed by the art in various manners which cover the stent with either a biologically active or an inactive covering which is less thrombogenic than the stent material and/or which has an increased capacity for promoting re-endothelialization of the stent situs. These solutions, however, all require the use of existing stents as substrates for surface derivatization or modification, and each of the solutions result in a biased or laminate structure built upon the stent substrate. These prior art coated stents are susceptible to delaminating and/or cracking of the coating when mechanical stresses of transluminal catheter delivery and/or radial expansion in vivo. Moreover, because these prior art stents employ coatings applied to stents fabricated in accordance with conventional stent formation techniques, e.g., cold-forming metals, the underlying stent substrate is characterized by uncontrolled heterogeneities on the surface thereof. Thus, coatings merely are laid upon the heterogeneous stent surface, and inherently conform to the topographical heterogeneities in the stent surface and mirror these heterogeneities at the blood contact surface of the resulting coating. This is conceptually similar to adding a coat of fresh paint over an old coating of blistered paint; the fresh coating will conform to the blistering and eventually, blister and delaminate from the underlying substrate. Thus, topographical heterogeneities are typically telegraphed through a surface coating. Chemical heterogeneities, on the other hand, may not be telegraphed through a surface coating but may be exposed due to cracking or peeling of the adherent layer, depending upon the particular chemical heterogeneity.